In radiography, stringent requirements are currently placed on the image quality of the x-ray images. In such images, as are taken especially in medical x-ray diagnosis, an object to be studied is exposed to x-radiation from an approximately point radiation source, and the attenuation distribution of the x-radiation is registered two-dimensionally on the opposite side of the object from the x-ray source. Line-by-line acquisition of the x-radiation attenuated by the object can also be carried out, for example in computer tomography systems.
Besides x-ray films and gas detectors, solid-state detectors are being used increasingly as x-ray detectors, these generally having a matrix shaped arrangement of optoelectronic semiconductor components as photoelectric receivers. Each pixel of the x-ray image should ideally correspond to the attenuation of the x-radiation by the object on a straight axis from the point x-ray source to the position on the detector surface corresponding to the pixel. X-rays which strike the x-ray detector from the point x-ray source in a straight line on this axis are referred to as primary beams.
The x-radiation emitted by the x-ray source, however, is scattered in the object owing to inevitable interactions, so that, in addition to the primary beams, the detector also receives scattered beams, so-called secondary beams. These scattered beams, which, depending on the properties of the object, can cause up to 90% or more of the total signal response of an x-ray detector in diagnostic images, constitute an additional noise source and therefore reduce the identifiability of fine contrast differences. This substantial disadvantage of scattered radiation is due to the fact that, owing to the quantum nature of the scattered radiation, a significant additional noise component is induced in the image recording.
In order to reduce the scattered radiation components striking the detectors, so-called antiscatter grids are therefore interposed between the object and the detector. Antiscatter grids consist of regularly arranged structures that absorb the x-radiation, between which transmission channels or transmission slits for minimally attenuated transmission of the primary radiation are formed. These transmission channels or transmission slits, in the case of focused antiscatter grids, are aligned with the focus of the x-ray tube according to the distance from the point x-ray source, that is to say the distance from the focus. In the case of unfocused antiscatter grids, the transmission channels or transmission slits are oriented perpendicularly to the surface of the antiscatter grid over its entire area. However, this leads to a significant loss of primary radiation at the edges of the image recording, since a sizeable part of the incident primary radiation strikes the absorbing regions of the antiscatter grid at these points.
In order to achieve a high image quality, very stringent requirements are placed on the properties of x-ray antiscatter grids. The scattered beams should, on the one hand, be absorbed as well as possible, while on the other hand, the highest possible proportion of primary radiation should be transmitted unattenuated through the antiscatter grid. It is possible to achieve a reduction of the scattered beam component striking the detector surface by a large ratio of the height of the antiscatter grid to the thickness or diameter of the transmission channels or transmission slits, that is to say by a high aspect ratio.
The thickness of the absorbing structure elements or wall elements lying between the transmission channels or transmission slits, however, can lead to image perturbations by absorption of part of the primary radiation. Specifically when solid-state detectors are used, inhomogeneities of the grids, that is to say deviations of the absorbing regions from their ideal position, cause image perturbations by projection of the grids in the x-ray image.
In order to minimize image perturbations due to antiscatter grids, it is known to move the grids in a lateral direction during the recording. In the case of very short exposure times of, for example, 1–3 ms, however, stripes may also occur in the image if the movement speed of the grids is insufficient. Even in the event of very long exposure times, perturbing stripes may occur owing to reversal of the grid movement direction during exposure.
In recording x-ray images, increasing use has recently been made of solid-state detectors which are formed from a plurality of an array of detector elements. The detector elements are arranged in this case in a generally square or rectangular grating. In the case of such solid-state detectors, as well, there is a need to employ effective suppression measures to reduce as far as possible the striking of scattered beams on the detector surface formed by the detector elements. Because of the regular structuring of the pixels, formed by the detector elements, of the detector, there is here, in addition, the risk of mutual interference between the structures of pixels and antiscatter grids. Disturbing moiré phenomena can thereby arise. These can certainly in specific instances be minimized or removed by a downstream image processing measure. However, this is possible only when their projection image on the detector is absolutely immutable.
The same problem occurs in nuclear medicine, especially when using gamma cameras, for example Anger cameras. With this recording technique also, as with x-ray diagnosis, it is necessary to ensure that the fewest possible scattered gamma quanta reach the detector. In contrast to x-ray diagnosis, the radiation source for the gamma quanta lies inside the object in the case of nuclear diagnosis. In this case, the patient is injected with a metabolic preparation labeled with particular unstable nuclides, which then becomes concentrated in a manner specific to the organ.
By detecting the decay quanta correspondingly emitted from the body, a picture of the organ is then obtained. The profile of the activity in the organ as a function of time permits conclusions about its function. In order to obtain an image of the body interior, a collimator that sets the projection direction of the image needs to be placed in front of the gamma detector. In terms of functionality and structure, such a collimator corresponds to the antiscatter grid in x-ray diagnosis. Only the gamma quanta dictated by the preferential direction of the collimator can pass through the collimator, and quanta incident obliquely to it are absorbed in the collimator walls. Because of the higher energy of gamma quanta compared with x-ray quanta, collimators need to be made many times higher than antiscatter grids for x-radiation.
For instance, scattered quanta may be deselected during the image recording by taking only quanta with a particular energy into account in the image. However, each detected scattered quantum entails a dead time in the gamma camera of, for example, one microsecond, during which no further events can be registered. Therefore, if a primary quantum arrives shortly after a scattered quantum has been registered, it cannot be registered and it is lost from the image.
Even if a scattered quantum coincides temporally—within certain limits—with a primary quantum, a similar effect arises. Since the evaluation electronics can then no longer separate the two events, too high an energy will be determined and the event will not be registered. Both said situations explain how highly effective scattered beam suppression leads to improved quantum efficiency in nuclear diagnosis as well. As the end result, an improved image quality is thereby achieved for equal dosing of the applied radionuclide or, for equal image quality, a lower radionuclide dose is made possible, so that the patient's beam exposure can be reduced and shorter image recording times can be achieved.
In future, increasing use will also be made for recording gamma images of solid-state detectors which are formed from an array of detector elements. The detector elements are arranged in this case in a generally square or rectangular grating. In the case of such solid-state detectors, as well, there is a need to employ effective suppression measures to reduce as far as possible the striking of scattered beams on the detector surface formed by the detector elements. Because of the regular structuring of the pixels, formed by the detector elements, of the detector, there is here, in addition, the risk of mutual interference between the structures of pixels and collimators.
Collimators for gamma cameras are generally produced from mechanically folded lead lamellae. This is a relatively cost-efficient solution. However, it has the disadvantage that, in particular when using solid-state cameras with an array of detector elements, for example in the case of cadmium-zinc telluride detectors, perturbing aliasing effects can arise because the structure of these collimators is then relatively coarse.
The publication by G. A. Kastis et al., “A Small-Animal Gamma-Ray Imager Using a CdZnTe Pixel Array and a High Resolution Parallel Hole Collimator” discloses a method for producing a cellularly constructed collimator for gamma radiation. In this case, the collimator is produced from laminated layers of metal films, here made of tungsten, which are photochemically etched. However, on account of the large number of photolithographic exposure and etching steps, this production method is very elaborate and cost-intensive.
U.S. Pat. No. 6,021,173 A describes an approach which is intended to avoid moiré structures during operation of an x-ray detector having an array of detector elements in conjunction with an antiscatter grid arranged in a stationary fashion. In this publication, the antiscatter grid is applied directly to the x-ray detector over the detector surface. The absorbing structure elements of the antiscatter grid are designed at a spacing from one another which is smaller than the extent of the smallest resolvable detail in the x-ray image. The regularly arranged absorbing structure elements are consequently formed at so high a spatial frequency
The post-published German patent application DE 101 51 568 discloses a method for applying an antiscatter grid to an x-ray detector in the case of which a basic structure for the antiscatter grid is produced directly on the detector surface by way of a rapid prototyping technique such that absorbing regions of the antiscatter grid are situated in less sensitive intermediate regions of the x-ray detector. However, the risk exists in this method of damaging the x-ray detector when producing the antiscatter grid.